Diabetes and Glycemic Control
Diabetes mellitus is a disease of major global importance, increasing in frequency at almost epidemic rates, such that the worldwide prevalence in 2006 was 170 million people and predicted to at least double over the next 10 to 15 years. Diabetes is characterized by a chronically raised concentration of glucose in the blood (hyperglycemia), due to a relative or absolute lack of the pancreatic hormone, insulin. Within the healthy pancreas, beta cells located in the Islets of Langerhans continuously produce and secrete insulin according to the glucose levels, maintaining near-constant glucose levels in the body.
Much of the burden of the disease to the patient and to health care resources is due to long-term tissue complications, which affect both the small blood vessels (microangiopathy, causing eye, kidney and nerve damage) and the large blood vessels (causing accelerated atherosclerosis, with increased rates of coronary heart disease, peripheral vascular disease and stroke). There is now good evidence that morbidity and mortality of diabetic patients is related to the duration and severity of hyperglycemia. (DCCT Trial, N Engl J Med 1993, 329: 977-986; UKPDS Trial, Lancet 1998; 352: 837-853; BMJ 1998; 317, (7160): 703-13; EDIC Trial, N Eng! J Med 2005, 353, (25): 2643-53).
In theory, returning glucose levels to normal by hormone replacement therapy using insulin injections and/or other treatments should prevent complications, but, frustratingly, near-normal glucose concentrations are very difficult to achieve and maintain in many patients, particularly those with type I diabetes. In these patients, glucose concentration can swing between very high (hyperglycemia) and very low (hypoglycemia) levels in an unpredictable manner. Thus, tight glycemic control is required. This control can be achieved by substituting the two functions of the normal pancreas—glucose monitoring and insulin delivery—for maintaining tight glycemic control. Furthermore, a closed loop or semi closed loop system provided with a feedback mechanism connecting between both functions (often referred to as an “artificial pancreas”) could theoretically maintain near-normal glucose levels.
Glucose Monitoring
Most diabetic patients currently measure their own glucose level periodically, i.e., several times during the day by obtaining finger-prick capillary samples and applying the blood to a reagent strip for analysis in a portable meter. Unfortunately, the discomfort involved leads to poor patient compliance. Testing cannot be performed while sleeping and while the subject is occupied. In addition, the results do not give information regarding the trends in glucose levels, but rather provide only discrete readings, taken at large time intervals from one another. Therefore, continuous glucose monitoring is advantageous, providing short-interval, essentially continuous glucose readings by performing discrete measurements, at a very high rate.
Glucose Monitoring Technologies
Continuous glucose monitoring can be performed by various methods and technologies, where most methods apply either non-invasive or minimally-invasive means.
Non-Invasive Continuous Glucose Monitoring
Non-invasive continuous glucose monitoring includes the sensing of glucose in blood, interstitial fluid (ISF) or other physiological fluids, primarily using optical means.
Continuous glucose monitoring based on optical methods employs various sensing methodologies for measuring glucose concentration levels. Optical sensing methods are quite prevalent among glucose sensors and include NIR, IR, Raman, Fluourescence, Polarimetry, and Photoacoustic (PA) technology.
In Near-Infrared (NIR) spectroscopy, a selected band of NIR light is transmitted through the sample, and the analyte concentration is obtained by the analysis of the resultant spectral information. The NIR absorbance bands tend to be broad and overlap, and are highly influenced by temperature, pH, and other physical factors. Nevertheless, the NIR spectrum allows for large optical path lengths to be used due to relatively easy passage through water (the light absorbance is directly proportional to the path length according to the Beer-Lambert law). U.S. Pat. No. 6,928,311 to Pawluczyk et al., assigned to NIR Diagnostics, Inc., describes a non-invasive monitor that uses NIR light. A beam of light in the NIR range is focused on the person's finger for about 30 seconds. By applying mathematical algorithms on the emerging light signal, the concentration of various blood analytes including glucose are determined and displayed to the user.
The NIR spectrum spans a wide range from 700 to 2500 nm. Absorption features throughout this spectral range primarily correspond to overtones and combinations of molecular vibrations. The absorption properties of water play a critical role in the regions of the NIR spectrum available for noninvasive measurements. Strong water absorption bands centered at approximately 1333, 1923, and 2778 nm (7500, 5200, and 3600 cm−1) create three transmission windows through aqueous solutions and living tissue. These spectral windows are termed the short-wavelength region (700-1370 nm, 286-7300 cm−1), the first overtone region (1538-1818 nm, 6500-5500 cm−1), and the combination region (2000-2500nm, 5000-4000 cm−1). Absorption features in the combination region correspond to first-order combination transitions associated with bending and stretching vibrations of C—H, N—H, and O—H functional groups. The first overtone region corresponds to the first-order overtone of C—H stretching vibrations, and the short-wavelength region includes numerous higher order combination and overtone transitions. For combination spectra, molar absorptivities are larger and bands are narrower compared to first overtone spectral features. NIR absorption features become significantly weaker and broader as the order increases, thereby greatly reducing the analytical utility of the short-wavelength region in terms of molecular vibrational information. (Anal Chem 2005 (77), pp. 5429-5439).
A relative dip in the water absorbance spectrum opens a unique window in the 2000-2500 nm wavelength region, saddled between two large water absorbance peaks. This window allows pathlengths or penetration depths on the order of millimeters and contains specific glucose peaks at 2130, 2270 and 2340 nm. This region offers the most promising results for quantifiable glucose measurements using NIR spectroscopy (Biomed Photonics Handbook, 2003, p. 18-13).
The different spectral regions permit for several sample volumes and optical path lengths: larger samples are possible for spectra collected at shorter wavelengths and longer wavelengths are restricted to smaller samples. Optimal sample thickness for the combination, first overtone, and short wavelength range are 1, 5, and 10 mm respectively. However, when the collected spectra encompass multiple spectral regions, it is not possible to match the sample thickness with each spectral region (Anal. Chem. 2005, 77, 5429-5439). Comparison between transmittance and reflectance measurements in glucose using near infrared spectroscopy shows that transmittance is preferred for glucose monitoring (Journal of Biomedical Optics 11 (1), pp. 014022-1-7, January/February 2006). FIG. 1a shows the optical absorption spectra of glucose in the NIR region for aqueous glucose after water subtraction (Journal of Biomedical Optics 5 (1), 5-16 Jan. 2000)
In mid-Infrared (mid-IR) spectroscopy, the wavelengths of glucose absorbance in the mid-IR spectrum range (2500-10000 nm) are used for the analysis of glucose concentration. Although the absorption bands tend to be sharp and specific, there is strong background absorption by water that severely limits the optical path length that may be used. FIG. 1b shows the optical absorption spectra of glucose in the mid-IR region for aqueous glucose after water subtraction (Journal of Biomedical Optics 5 (1), 5-16 Jan. 2000).
In Raman spectroscopy, Raman spectra are observed when incident light is inelastically scattered producing Stokes and anti-Stokes shifts, where the latter is the more prevalent. Raman spectra are less influenced by water compared to NIR/IR and the peaks are spectrally narrow. In addition, Raman spectroscopy requires minimal sample preparation. However, the signal is weak and therefore requires a highly sensitive detection system (e.g., CCD array).
It is possible to detect glucose by monitoring the 3448 nm (2900 cm−1) C—H stretch band or the C—O and C—C stretch Raman bands at 8333-11111 nm (900-1200 cm−1), which represents a fingerprint for glucose (Clinical Chemistry 45:2 165-177, 1999). FIG. 1c shows the Raman spectrum for aqueous glucose, after subtraction of the water background (Journal of Biomedical Optics 5 (1), 5-16 Jan. 2000).
In fluorescence energy transfer (FRET)-based assay for glucose measurement, concanavalin A is labeled with the highly NIR-fluorescent protein allophycocyanin as donor and dextran labeled with malachite green as the acceptor (see, J Photochem Photobiol 2000; 54: 26-34. and Anal Biochem 2001; 292: 216-221). Competitive displacement of the dextran from binding to the lectin occurs when there are increasing glucose concentrations, leading to a reduction in FRET, measured as intensity or lifetime (time-correlated single-photon counting).
Polarimetry involves the optical rotation of the polarized light by the chiral centers of glucose, which is determined by the structure of the molecule, the concentration of the molecule, and the optical path length the light traverses through the sample. Each optically active substance has its own specific rotation, as defined by Biot's law. The measurement of the optical rotation requires a very sensitive polarimeter, due to the low glucose concentrations in the cell. For example, at a wavelength of 670 nm, glucose will rotate the linear polarization of a light beam approximately 0.4 millidegrees per 10 mg/dl for a 1-cm sample pathlength (Biomedical Photonics Handbook, 2003, p. 18-14). In addition, the presence of other optically active molecules makes the accurate detection of glucose concentration complicated.
Finally, PA spectroscopy involves light which is absorbed by glucose, leading to thermal expansion and to the generation of a detectable ultrasound pressure wave. In one study, solutions of different glucose concentrations were excited by NIR laser pulses at wavelengths that corresponded to NIR absorption of glucose in the 1000-1800 nm range. There was a linear relationship between PA signal and glucose concentrations in aqueous solutions (Diabetes Technology & Therapeutics, Vol. 6, November 2004, O. S. Khalil). This method is particularly sensitive to changes in temperature.
Optical glucose measurement techniques are particularly attractive for several reasons: they utilize nonionizing radiation to interrogate the sample, are reagentless, and fast. The use of optical glucose monitoring methods is especially attractive because they are nondestructive and reagentless, thereby eliminating the risk of unsafe reactions and their by-products.
Although optical approaches for glucose sensing are attractive, they are nevertheless often plagued by a lack of sensitivity and/or specificity since variations in optical measurements depend on variations of many factors in addition to glucose concentration. Isolating those changes which are due to glucose alone and using them to predict glucose concentration is a significant challenge in itself (Journal of Biomedical Optics 5 (1), 5-16 Jan. 2000). Furthermore, non-invasive optical glucose monitors, which involve sensing of glucose levels through the skin, involve very low signal-to-noise ratio, scattering and interferences by bodily fluids and by the skin itself, causing noninvasive optical sensors to lack specificity and repeatability.
Since optics-based noninvasive applications do not produce accurate and specific results, it would be desirable to provide an immediate application of the optical methods directly to the ISF or to fluids comprising endogenous components of the ISF, thus, eliminating the attenuating effects of the skin.
Invasive Continuous Glucose Monitoring
Invasive continuous glucose monitoring involves the implantation of a sensing device in the body. As detailed in U.S. Pat. Nos. 6,122,536 to Sun et al. and 6,049,727 to Crothall, assigned to Animas Corporation, an invasive spectroscopy-based glucose sensor, designed for long-term (>5 years) internal use is under development. The Animas sensor has the advantage of being able to directly read glucose in the blood. A small, ultralight C-clamp detector is surgically implanted around a 4-5 mm (0.2 inch) diameter blood vessel. The detector has two tiny probes at the tips of the C-clamp structure which puncture each side of the vessel and allow transmission of a clean infrared light signal between them. A larger device housing a laser generator plus signal analysis is located nearby within a closed compartment under the skin. The laser IR signal is transmitted to the detector around the vessel and returns the transmitted beam back to the processing unit. Readings are available at short time intervals. Major advantages of the implantable approach are that calibration is required only once a week and that although minor surgery is required, this sensor provides direct access to blood.
Other invasive continuous glucose monitoring (CGM) systems are often based on electrochemical techniques. U.S. Pat. No. 6,862,465 to Shults et al. and U.S. Pub. No. 2006/0036145A1 to Brister et al., assigned to DexCom, Inc., describe a long term glucose oxidase (GOX)-based CGM system. The system includes a sensor, a small implantable device that continuously measures glucose levels in the subcutaneous tissue, and a small external receiver to which the sensor transmits glucose levels at specified intervals. The receiver displays the patient's current blood glucose value, as well as 1-hour, 3-hour and 9-hour trends. The receiver also sounds an alert when an inappropriately high or low glucose excursion is detected. The DexCom™ Long Term Sensor is implanted under the skin in the abdomen by a local anesthetic short procedure carried out by a physician. This sensor is designed to function for up to one year. At the end of its life, the sensor can be removed by a physician in a short procedure, and another sensor implanted.
Implanted CGMs have several disadvantages, such as the need for a surgical procedure to implant the device, possibility of failure, the potential for sensor blockage, and biocompatibility problems.
Minimally-Invasive Glucose Monitoring
Disadvantages in the performance and operation of non-invasive and fully invasive CGMs lead to the development of various minimally-invasive CGM systems. Minimally-invasive CGMs often measure glucose levels in the ISF within the subcutaneous tissue, and are based on various sensing technologies. The strong correlation between blood and ISF glucose levels, allows for accurate ISF glucose measurements (Diabetologia 1992; 35, (12): 1177-1180).
GlucoWatch® G2® Biographer is one commercially available minimally-invasive glucose monitor. GlucoWatch® is based on reverse iontophoresis as detailed in U.S. Pat. No. 6,391,643, to Chen et al., assigned to Cygnus, Inc. A small current passed between two skin-surface electrodes draws ions and (by electro-endosmosis) glucose-containing ISF to the surface and into hydrogel pads provided with a GOX biosensor (JAMA 1999; 282: 1839-1844). Readings are taken every 10 minutes with a single capillary blood calibration.
Disadvantages of the GlucoWatch® include occasional sensor values differing markedly from blood values, skin rashes and irritation in those locations which are immediately underneath the device appearing in many users, a long warm up time of 3 hours, and skips in measurements due to sweating.
Two additional commercially available minimally-invasive monitors are GOX-based CGMs, based on enzyme-immobilization.
The Guardian® RT Continuous Glucose Monitoring System, developed by Medtronic MiniMed Inc. is a GOX-based sensor, as described in U.S. Pat. No. 6,892,085 to McIvor et al. The sensor consists of a subcutaneously implanted, needle-type, amperometric enzyme electrode, coupled with a portable logger (Diab Tech Ther 2000; 2: Supp. 1, 13-18). The Guardian® RT system displays updated glucose readings every five minutes, together with hypo- and hyperglycemic alarms. The sensor is based on the long-established technology of GOX immobilized at a positively charged base electrode, with electrochemical detection of hydrogen peroxide production.
U.S. Pat. No. 6,862,465 to Shults et al. and U.S. Pub. No. 2006/0036145A1 to Brister et al., assigned to DexCom, Inc., describe a short-term GOX-based CGM system. The system includes a sensor, a small insertable or implantable device that continuously measures glucose levels in the subcutaneous tissue, and a small external receiver to which the sensor transmits glucose levels at specified intervals. The receiver displays the patient's current blood glucose value, as well as 1-hour, 3-hour and 9-hour trends. The receiver also sounds an alert when an inappropriately high or low glucose excursion is detected. The DexCom™ STS™ Continuous Glucose Monitoring System is a user insertable short-term sensor that is inserted just under the skin where it is held in place by an adhesive. Once inserted the user would wear the sensor for up to three or seven days before being replaced. After three or seven days, the user removes the sensor from the skin and discards it. A new sensor can then be used with the same receiver. The DexCom™ STS™ Continuous Glucose Monitoring System has been FDA-approved.
The Freestyle Navigator™ is another GOX-based sensor, detailed in U.S. Pat. No. 6,881,551 to Heller et al., assigned to Abbott Laboratories, formerly TheraSense, Inc. This sensor is placed just under the skin by a disposable self-insertion device. Information is communicated wirelessly between the transmitter and the receiver every minute. The receiver is designed to display glucose values, directional glucose trend arrows, and rate of change. The receiver also has high and low glucose alarms, and stores glucose data for future analysis.
Numerous disadvantages inherent to glucose monitoring are present in CGMs which employ GOX-based reactions. Most GOX-based devices rely on the use of oxygen as the physiological electron acceptor, and thus, are subject to errors due to fluctuations in the oxygen tension and the stoichiometric limitation of oxygen in vivo. The amperometric measurement of hydrogen peroxide requires application of a potential at which additional electroactive species exist, e.g. ascorbic and uric acids or acetaminophen. These and other oxidizable constituents of biological fluids can compromise the selectivity and hence the overall accuracy of the glucose concentration measurement. Hydrogen peroxide deactivates the GOX molecules, limiting the time available for application of the sensor. Miniaturizing the sensing technology within the cannula, which requires high levels of enzyme loading, while keeping high measurement sensitivity, remains a challenge.
Numerous disadvantages inherent to glucose monitoring are present in CGMs which employ GOX-based reactions. Most GOX-based devices rely on the use of oxygen as the physiological electron acceptor, and thus, are subject to errors due to fluctuations in the oxygen tension and the stoichiometric limitation of oxygen in vivo. The amperometric measurement of hydrogen peroxide requires application of a potential at which additional electroactive species exist, e.g. ascorbic and uric acids or acetaminophen. These and other oxidizable constituents of biological fluids can compromise the selectivity and hence the overall accuracy of the glucose concentration measurement. Hydrogen peroxide deactivates the GOX molecules, limiting the time available for application of the sensor. Miniaturizing the sensing technology within the cannula, which requires high levels of enzyme loading, while keeping high measurement sensitivity, remains a challenge.
Microdialysis Based Glucose Monitors
Microdialysis is an additional commercially-available minimally-invasive technology (Diab Care 2002; 25: 347-352) for glucose monitoring as detailed in U.S. Pat. No. 6,091,976 to Pfeiffer et al., assigned to Roche Diagnostics GmbH, and the marketed device, GlucoDay® S, produced A. Menarini Diagnostics. A fine, semi-permeable hollow dialysis fiber is implanted in the subcutaneous tissue and perfused with isotonic fluid. Glucose diffuses across the semi-permeable fiber and is pumped outside the body via the microdialysis mechanism for measurement by a glucose oxidase-based electrochemical sensor. Initial reports (Diab Care 2002; 25: 347-352) show good agreement between sensor and blood glucose readings, and good stability with a one-point calibration over one day. Higher accuracies were found when using the microdialysis-based sensor, compared to the needle-type sensor (Diabetes Care 2005; 28, (12): 2871-6).
Disadvantages of the microdialysis-based glucose sensors stem primarily from the constant perfusion of solution through the microdialysis probe. This operational method requires the presence of a dedicated pump and reservoir, leading to large and bulky devices, and also necessitates high-energy consumption. Furthermore, the relatively large size of the microdialysis catheter often causes a wound and subsequent local tissue reactions, following its insertion into the subcutaneous tissue. Finally, the microdialysis process generates long measurement lag times, due to the essential slow perfusion rates and long tubing.
Optics Based Glucose Monitors
U.S. Pub. No. 2007/0004974 A1 to Nagar et al. describes a device for assaying an analyte in the body, comprising a light source implanted in the body, able to illuminate a tissue region with light, at a wavelength that is absorbed by the analyte and as a result generates PA waves in the tissue region. An acoustic sensing transducer is coupled to the body, receives acoustic energy from the PA waves, and generates responsive signals. A processor receives the signals and processes them to determine a concentration of the analyte in the illuminated tissue region.
U.S. Pat. No. 5,605,152 to Slate et al. describes an improved glucose sensor adapted for in vivo implantation which includes one or more optical fiber optrodes mounted within a semi-permeable probe housing designed for differential diffusion of glucose and oxygen. An enzyme optrode comprises an optical fiber with an enzyme coating such as GOX to yield an enzymatic glucose reaction. An oxygen sensitive coating such as a fluorescent dye is provided on the enzyme optrode close to the enzymatic reaction and also on a reference optrode at a position spaced substantially from the enzymatic reaction. Optical monitoring of the fluorescent activity of the optrode coatings provides an indication of oxygen depletion as a result of the enzymatic reaction and thus indicates the glucose concentration level. The semi-permeable housing is designed to ensure that the reaction proceeds with a stoichiometric excess of oxygen.
Disadvantages of optics-based glucose monitoring techniques stem primarily from the indirect optical means applied for measuring analyte concentration levels. In addition, optics-based techniques involving an electrochemical reaction possess the disadvantages inherent to glucose monitors which employ GOX-based reactions, described hereinbefore.